Methods of minimizing scattering and improving tissue sampling in non-invasive testing and imaging

ABSTRACT

An improved method and apparatus for use in optical testing of concentration in samples has been developed. The apparatus restricts the solid angle of illumination and the solid angle of detection to eliminate a high proportion of the scattered radiation while allowing the ballistic radiation and the snake-like radiation to be transmitted. In samples which contain multiple scattering centers, this allows less correction for variations in effective pathlength and allows easier calibration of the apparatus. The use of polarized radiation as a means of minimizing scattered radiation in the sample is also disclosed.

REFERENCE TO RELATED APPLICATIONS

This application is a divisional application of Ser. No. 09/406,675filed on Sep. 27, 1999 now abandoned, Pending, which is a divisionalapplication of Ser. No. 08/937,934 filed on Sep. 25, 1997, now U.S. Pat.No. 6,064,065, which in turn is a divisional application of Ser. No.08/479,955 filed on Jun. 7, 1995, now U.S. Pat. No. 5,672,875; which isa CIP of Ser. No. 08/333,758, Filed Nov. 3, 1994, now U.S. Pat. No.5,818,048, which is a CIP of Ser. No. 08/182,572, filed Jan. 14, 1994,now U.S. Pat. No. 5,424,545; which is a CIP of Ser. No. 08/130,257,filed Oct. 1, 1993, now U.S. Pat. No. 5,434,412; which is a CIP of Ser.No. 07/914, 926, filed Jul. 17, 1992, now U.S. Pat. No. 5,334,287. Thecontents of all of the aforementioned application(s) are herebyincorporated by reference.

BACKGROUND OF THE INVENTION

The present invention relates to improvements in optical systems andtheir uses for the measurement of concentration and temperature inscattering media, and the related discrimination of subsurface features.More particularly, the invention provides methods and apparatus whichminimize the ratio of diffusely scattered radiation to directlytransmitted radiation reaching the detector(s) in optical concentrationmeasurement and imaging apparatus. The methods and apparatus of theinvention have special applicability to noninvasive testing,particularly for concentration measurements of materials such as glucoseand hemoglobin in blood.

Recent literature is replete with articles describing attempts atperforming non-invasive testing using optical measurements (e.g.,infrared systems). Part of this expansion has been fueled by the spreadof acquired immunodeficiency disease syndrome (AIDS), and the associatedfear among public and health care personnel of AIDS. AIDS and otherdiseases such as hepatitis are born in the blood and can be spread byimproper practice of invasive procedures. In addition, the diabeticpopulation has also been anxiously awaiting non-invasive testinstruments for many years. Many diabetics must test their blood glucoselevels four or more time a day. The modern battery powered instrumentsfor home use require a finger prick to obtain the sample. The extractedblood samples are placed on a chemically-treated carrier which isinserted into the instrument to obtain a glucose reading. This fingerprick is painful and can be objectionable when required frequently. Inaddition, although the price has dropped considerably on theseinstruments, the cost for the disposable and the discomfort and healthrisk associated with having open bleeding is undesirable.

Accordingly, a number of groups have recently tried to make non-invasiveinstruments for testing a variety of analytes, particularly glucose. Arecent trend in non-invasive testing has been to explore the use of thenear infrared spectral region, primarily 700-1100 nm because this is thespectral response range of the silicon detectors typically used in theprior art. A wider wavelength range to ˜1800 nm can be accessed by theaddition of germanium and/or InGaAs detectors, and useful measurementscan be made into the 2500 nm range with InSb or other detectors. Theregion below ˜1400 nm is the most useful in transmission, as tissue istransparent enough there to allow high enough photon flux for accuratedetection. Above 1400 nm, the strong absorption of water limits thepenetration depth of tissue, so that useful measurements are typicallymade in reflectance geometry. Below 1100 nm, the penetration of thelight is sufficient that the signal modulation during the arterial pulsecan be measured comfortably in both transmission and reflectancegeometries. Above 1400 nm, such pulsatile measurements are extremelydifficult in transmission due to low intensity, and similarly difficultin reflectance because the light does not penetrate deeply enough tosample the pulsatile capillary beds.

Most of the non-invasive testing work has been carried out using classicspectrophotometric methods, such as a set of narrow wavelengths sources,or scanning spectrophotometers which scan wavelength by wavelengthacross a broad spectrum. The data obtained from these methods arespectra which then require substantial data processing to eliminatebackground; accordingly, the papers are replete with data analysistechniques utilized to glean the pertinent information. Examples of thistype of testing includes the work by Clarke, see U.S. Pat. No.5,054,487; and primarily the work by Rosenthal et al., see e.g., U.S.Pat. No. 5,028,787. Although the Clarke work uses reflectance spectraand the Rosenthal work uses primarily transmission spectra, both rely onobtaining near infrared spectrophotometric data.

The major successful application of non-invasive testing is themeasurement of hemoglobin oxygen saturation with pulse oximetry. Themost common method compares the percentage modulation of the intensityof light traversing a body part at two wavelengths chosen so that theratio of their respective modulations is a relatively strong function ofoxygen saturation. The observed change in this ratio is relatively largebecause the two hemoglobin species involved have both high enoughconcentrations and specific absorptions that they dominate the creationof the pulsatile signal components at the wavelengths of interest. As aresult, the ratio of modulations can be attributed substantially to thetwo hemoglobins alone, and only needs to be measured to the order of0.1% in order to achieve clinically significant detection limits withacceptable universality of calibration.

The optical system in typical pulse oximeters have two or more LEDemitters placed side-by-side on one side of a finger, and a singledetector receiving the radiation on the other side of the finger. Somemore recent systems have the detector on the same side of the tissue asthe emitters, with baffles preventing the direct illumination of thedetector by the sources. As the sources are physically small andoptically displaced from each other and the detector, the light fromeach detector enters the tissue at slightly different locations, andtherefore travel different paths through tissue to the detector.

Despite its relatively large signal levels, pulse oximetry haswell-known difficulties such as the selection of an adequately vascularsampling site on each individual and variability of the results withmotion of the site and breathing by the patient, as well as sensitivityto changes in blood pressure, heart rate, temperature, and tissuehydration. Disturbances such as motion and breathing artifacts typicallyappear as statistical outriders, i.e., as measurements which fall welloff the “average” calibration curve of the instrument obtained from agroup of individuals breathing controlled gas mixtures to vary theiroxygen saturation.

The calibration of a pulse oximeter is subject to these same errorsources; it is not uncommon to find site-to-site variations on the sameindividual, with results that suggest that the calibration curve evenvaries, for example, with the absolute magnitude of the pulsatile signalmodulation. The effort to obtain a meaningful universal calibration isclearly at odds with intra- and inter-individual physiologicalvariations.

Despite recent efforts to improve the measurement S/N by increasingsource intensities and lowering detector noise, as well as increasingthe number of detectors, the frequency of outriders and the universalityof calibration have not improved substantially. Thus it is clear thatwhile the light traversing the tissue is being measured more precisely,the site- and physiologically-induced variability has not been improvedsignificantly below the 0.1% level needed for the measurement of oxygensaturation.

While these physical and physiological interferences are marginallyacceptable for oxygen saturation measurements, they set a lower limit ofdetectivity that is too high for other clinical analytes such as glucoseand cholesterol for which the combination of concentration and specificabsorption requires optical measurements to be made 100-1000 times moreprecise than for the hemoglobins used in pulse oximetry. Thehemoglobins, which in themselves are difficult to calibrate in thepresence of these site- and physiologically specific limitationscomprise a major background interference for the measurement of suchtrace constituents as glucose.

The optical systems employed for these lower concentration analytesnaturally drew on the experience of pulse oximetry, and typically employsimilar arrangements of a plurality of slightly displaced LED's toextend the wavelengths sampled, or which use fiber optics to carry lightto and from the sources and/or spectrometers which perform theseparation of the signal into the different wavelengths employed.Displacement of the sources and wide numerical apertures for the lightentering and leaving the tissue enhance the likelihood that differentdetected wavelengths will have sampled different portions of the medium.Many of the physiological interferences to accurate measurement aremediated by differences in the mean paths traced by light of differentwavelength in traversing the intervening tissue between light source[s]and detector[s]. These path variations are produced, in part, by lightscattering in the tissue, which varies with the wavelength of the lightand which makes photons follow a jagged overall path from scattering toscattering. The detected signals are a complicated function of both thescattering and the total absorption of all constituents along the longertotal path of the light. Thus, the present optical systems used fornon-invasive measurement allow and perhaps even encourage light ofdifferent wavelengths to travel different paths through the tissue,sampling lateral and axial tissue inhomogeneities differently.

This situation violates a fundamental premise of all opticalnon-invasive measurement methods; namely, that the light intensity whichis measured in the individual detection channels can be attributed tothe analyte and not to any difference in tissue sampling. Tissueinhomogeneity produces wavelength-dependent spreading of the light whichultimately reaches the detectors, and in the extreme of high scatteringand large inhomogeneity, the mixture of detector signals becomes anuncontrollable and uncalibratable average response to the physiologicaland biochemical conditions at the sampled site.

In addition the existing non-invasive art has employedspectrophotometric methods which limit the intensity of light detectedin the individual resolution elements, and which also apply the methodin a way which uses the available spectral information inefficiently.These methods were conceived primarily for accurate determination ofnarrow band spectral structures rather than for discriminating thepresence of weak broadband features in strong broadband backgrounds thatcharacterize the non-invasive measurement problem for constituents suchas glucose. The multivariate analysis mathematics required to separatethe analyte signature from strongly overlapping interferent signaturesalso introduces an error propagation penalty that compounds theintensity limitation by increasing the impact of detector noise on thecalculated measure of concentration.

Improvements that enhance the solution of problems of interference inbroadband spectra, by obtaining different raw data, are described inU.S. Pat. No. 5,321,265 (the “Block '265 Patent”). This patent setsforth a different approach in non-invasive testing as compared with theprior instruments and methods. As noted, substantially all workers inthe non-invasive testing field prior to the Block '265 Patent were usingclassic spectrophotometric instrumentation and substantial processing inan attempt to resolve the low resolution features from the background.However, the spectra of analytes such as glucose in a human body are notdiscrete high resolution features which spectrophotometric instrumentswere originally designed to measure but rather have a few lowresolutions features with much of the information contained in subtlevariations of the detected intensity as a function of wavelength. Assuch, these spectra appear more like reflectance spectra of coloredobjects in the visible region. The Block '265 Patent teaches the use ofan analog of human color perception to obtain meaningful data by meansof methods and apparatus which utilize overlapping, broadbeam detectorsto mimic the spectral response characteristics of the human retinalcones, but translated into the near-infrared. This approach, which isradically different than classic spectrophotometric measurements,provides advantageous effects in determining the concentration ofglucose and other similar materials in an aqueous solution and isparticularly advantageous for use with scattering media such as tissuewhere it also provide the added advantage of higher light flux at thedetectors so that the intrinsic shotnoise limitation as a percentage ofthe total signal intensities is reduced.

U.S. patent applications Ser. No. 08/130,257 now U.S. Pat. No.5,434,412; Ser. No. 08/182,572 now U.S. Pat. No. 5,424,545 and Ser. No.08/333,758 now U.S. Pat. No. 5,818,048, the disclosures of which areincorporated herein by reference, all disclose improvements in the basictechniques and apparatus described in the Block '265 Patent. Theseimprovements include the concepts of congruent illumination anddetection of light emerging from the sampled tissue site, pulsatileprocessing, modulation of illumination sources as a means of eliminatingunwanted radiation, the use of non-overlapping broad beam radiation aswell as overlapping radiation, and a number of variations thereon. Theseapplications make it clear, in part, that a variety of techniques areuseful (and in some instances may be necessary) to deal with theproblems encountered in non-invasive measurement of analyteconcentration in tissue or other scattering media. Many of theseproblems arise from the fact that scattering media exhibit highereffective path lengths than their physical dimensions because ofscattering by the samples themselves. In fact, the samples, such ashuman tissue, act as if they are formed of a plurality of scatteringsites or centers in the sample. Techniques such as the congruentillumination and congruent detection described in these patentapplications equalize the acceptance angles and distances traveled bylight of different wavelengths outside the scattering media.Technically, this is achieved by locating all the illumination sourcesand/or detectors so that the path lengths and angles between the mediaand the detector(s)/source(s) are equal, so that the detectors orradiation sources act as if they were optically superimposed.

However, the desired congruency of detected light is degraded within theobserved media because the multiple scatterings of light spread thelight beam to adjacent regions in a way which is strongly wavelengthdependent. If the scattering media is inhomogeneous, the result of thisspreading is to mix light from these adjacent structures in relativeamounts which are dependent on wavelength. One object of the presentinvention is to reduce this disturbing effect by refining the launch anddetection optics to limit their angular acceptance ranges.

It has long been known that a certain portion of the illuminatingradiation survives transit across a turbid sample without being eitherscattered or absorbed, while a much larger portion is scattered in alldirections. The more scattering a particular photon undergoes the longerthe integrated path it follows, and the longer the time that elapsesbefore it emerges from of the sample. Some groups have attempted toreduce the deleterious effects of scattering by using pulsed sources andtime gating the detection so as to view the sample only in light whichhas undergone few scatterings. What is measured is a “snapshot” of thesample in light starting at the time of flight for an unscattered beam,and extending long enough in time to obtain sufficient signal for thedesired analysis without including much scattered radiation. When thetime gate is short, “ballistic” or “snakelike” photons which haveundergone no or few scatterings along their path are selected, andshadowgram images similar to those commonly obtained with x-ray's can beobtained.

This approach, however, requires complicated apparatus, and in additionto the intensity limitation from the short time-gate after each pulse ofthe light source, adds a further limitation on the number of detectedphotons because the duty-cycle of the pulsed source is low compared tothe continuous source of the present invention. Other research groupssuch as Wist et al., IEEE Transactions on Medical Imaging, 12 (4)751-757 (1993), have demonstrated that shadowgram-type images can beobtained by severely restricting the angular acceptance range ofdetected photons about the forward direction, essentially demonstratingthat doing so limits the detection to “ballistic” or “snakelike”photons. The Wist et al. apparatus generates a geometrically narrow beamwhich is raster-scanned across the sample, at a first wavelength, andthen generates new images at changed wavelengths. The work of thisgroup, however, also demonstrates a severe limitation on the total fluxof transmitted photons which make it inapplicable to the detection oftrace constituents in scattering media.

Other workers such as Schmitt et. al., SPIE 1641, 150-161, (1992), havedemonstrated advantages for using collimated input and output light onin vitro phantoms that simulate some of the light scattering propertiesof turbid media, but the transmitted intensity limitation of theirsystem when it was applied to real in vivo measurements made itnecessary to change the system design away from this collimatedapproach. One difficulty appears to be that their in vitro system wasdesigned to “approximate the plane-parallel conditions under which [thetheoretical] photon diffusion model was derived,” rather than addressingthe characteristics of the in vivo sample. Schmitt's collimated systemwas designed to approximate a “collimated beam of infinite extent” byestablishing a finite incident beam of light traversing tissues andconfining the collimated detection to a small central region on the exitside, apparently in order to eliminate unwanted edge effects. Inaddition, the narrowband sources and detector used limited thetransmitted intensity.

The failure of Schmitt's design was that insufficient photon flux wasavailable at the detector, so that this system was abandoned for his invivo work. Instead, Schmitt's in vivo apparatus employed a fiber opticthat launched light into the tissue at its large (˜50 degree) numericalaperture, and an integrating detector on the opposite side of the tissuereceiving light through almost the whole hemisphere. Even then, as notedin his article, the system had inadequate light intensity for themeasurement he was attempting. His work thus vividly illustrates thelight transmission limitations of real tissue that characterizes theprior art.

Thus, it is a specific object of the present invention to balance thelight collection efficiency and spatial resolution of the opticalsampling system viewing scattering media to simultaneously achieve highdetected light intensity and equality of response, as a function ofwavelength, to inhomogeneous inclusions within the media This isaccomplished by selecting optical configurations of sources, detectors,and intervening optical elements to minimize the effect of tissueinhomogeneities on the relative changes in signal strengths in each ofthe different detectors due to the presence of analyte.

It is a further object of the invention to achieve this balance in a waywhich improves the repeatability of the measurements from site-to-siteon a given individual in the presence of disturbances such as motion,breathing, hydration, and the like, with the ultimate objective toachieve universal calibratability of the measurement across subject inthe presence of such disturbances.

A related object of the invention is to provide a method of non-invasiveconcentration measurement in a scattering media which increases theratio of direct collimated radiation to diffusely scattered radiationreaching the detector, while maintained high integrated light intensityat the detectors.

Another object in the invention is provide an apparatus for non-invasiveconcentration measurements which maximizes the ratio of directcollimated radiation to diffusely scattered radiation while maintaininghigh integrated light intensity.

A further object of the invention is to facilitate the use of tightercollimation by increasing the number of photons received in theindividual detector resolution elements through broadening theirwavelength acceptance range.

A similar objective of the invention is to facilitate the use of tightercollimation by increasing the number of photons received by individualdetector resolution elements through increasing their surface area whilemaintaining their congruency.

Yet another object of the invention is to further facilitate the use oftighter collimation by the use of overlapping broadband detectorresolution elements in an analog of human color perception to combineincreased photon flux with more efficient separation of similar broadanalyte and interferent spectral features.

Consequently, it is a specific object of this invention to selectoptical configurations of sources, detectors, and intervening opticalelements to minimize the effect of tissue inhomogeneities on therelative changes in signal strengths in each of the different detectorsdue to the presence of analyte.

It is a still further object of this invention to adjust the opticalinterface to take maximum advantage of the natural spreadingcharacteristics of the light distribution patterns in tissue inmaximizing the S/N of the determination.

These and other objects which features the invention will be apparentfrom the detailed description and the drawing.

SUMMARY OF THE INVENTION

The present invention features methods and apparatus for measuringconcentration in a sample which contains a plurality of radiationscattering sites, and for measuring the distribution of concentrationand/or temperature within a sample when employed with imaging detectors.The methods and apparatus can also be utilized for discrimination ofsubsurface features through shadowgram generation. This procedure isalso useful for detection of temperature inhomogenieties. The methodsand apparatus of the invention employ means for restricting the solidangle of illumination and/or collection, e.g., by collimation of theradiation to minimize the amount of scattered radiation collected,employ polarization of the illuminating radiation to differentiatescattered from unscattered radiation, or a combination of the two. Themethods and apparatus of the invention provide more reproduciblemeasurements on scattering media and are particularly well suited tonon-invasive testing of tissue for materials such as glucose and drugsof abuse.

More particularly, the present invention provides a method of measuringthe concentration in a sample of a selected substance which absorbsradiation in a particular region of the spectrum. The sample containingthe substance of interest also contains a plurality of sites whichscatter radiation in the same particular region of the spectrum. Themethod has the steps of illuminating the sample with broad geometricarea illumination within a particular region of the spectrum (preferablyusing broad spectrum radiation) where the substance of interest hasabsorption, with the illumination and detection solid angles restricted,and with both said solid angles extending over a geometrically widesurface cross-section. The term “broad spectrum illumination” as usedherein means and implies that the wavelength of the illumination coversa substantial portion of the region of the spectrum in which there isabsorption by the selected substance. Normally, the illumination isgreater than 50 nm wide, and if the substance of interest has absorbanceat several wavelengths in the particular region, it preferably is wideenough to cover all of the absorption bands.

After leaving the sample, radiation which is transmitted or reflectedfrom the sample is collected with a detector, the detector beingselected and located such that the each resolution element of thedetector collects radiation only from a limited solid angle extendingover a relatively wide area of the viewed surface. The restricted solidangle illumination also extends over a relatively wide area of theilluminated surface. The term “wide” as used herein implies a beam widththat is comparable or larger in size than the thickness of the tissuebeing viewed. The term “comparable” means and implies that the width ofthe beam or viewing area is at least half the thickness of the sample ortissue. That thickness is itself preferably restricted to be not muchdeeper than several “natural” 1/e penetration lengths, the depth overwhich the diffuse radiation photon density falls to a few percent of itsmaximum value near the entrance surface of the medium. For typicaltissues viewed in the 700-1400 nm wavelength range, these preferredthicknesses are of the order of a few mm, and the surface areas throughwhich the light enters and exits are both in the 5-10 mm range.

The terms “restricted solid angle” and “limited solid angle” as usedherein imply that the type of detector or illuminating radiation, whichmay include some form of filtering and/beam focusing apparatus, limitsthe angle over which the illumination or detection occurs. Preferably,the illumination and or detection is restricted to a solid angle ofabout 10° or less from the central illumination beam axis. Thisconfiguration maximizes-the ratio of directly transmitted radiation toscattered radiation collected by the detector from the sample.Preferably, the restricted solid angle of illumination is achieved bycollimating the radiation from the radiation source prior toilluminating the sample, most preferably with collimating optics such asdescribed herein. Alternatively, a laser or another source whichprovides restricted divergence illumination may be used without thenecessity of some type of collimating optics.

Similarly, the preferred detector limits the solid angle of the detectedradiation transmitted from said sample by excluding uncollimatedradiation prior to collection by the detector. Again, collimatingoptics, such as a combination of lenses and/or apertures, can be used.Alternative collimating optics known in the art, such as a channel plateor a honeycomb collimator, can be used as well. Some of the benefits ofthe method may be achieved by comparing the results for two or morenarrow angular acceptance ranges at different angles relative to thecentral axis of the illumination beam. In fact, using a second, off axisdetector can assist in identifying the contribution of diffuse radiationand assist in correction for motion and other artifacts. In anotherembodiment, it is advantageous to have congruent detectors viewing thescattering media on the same side as the illumination, in what iscommonly referred to as a reflectance or transflectance measurementgeometry. Certain tissue sites, such as the forehead, may beparticularly advantageous in this geometry because the vascular tissuebetween the skin and the bone has a thickness comparable to the“natural” penetration depths, and the bone serves as a relatively inertbacking that isolates the overlying tissue from other analyte containingdeeper tissue.

In addition to the strictly collimated beam, a combination of selectedslightly converging or diverging beams generated by a combination oflenses and/or apertures can be used in the present invention. In oneaspect of the invention, a first converging lens is selected and locatedon the illumination side of the sample such that its focal point for theilluminating radiation is located on the detection side of the sampleand the first converging lens. In this embodiment, limiting the solidangle viewed by the detectors may be achieved by means of a secondconverging lens between the sample and the detector which has a limitingaperture mask at its focal distance so that the size of a central holein that mask defines the angular acceptance range. This same type ofsecond converging lens is used on the illumination side. After passingthrough this aperture, the light beam is then expanded once again andpassed through the congruent beamsplitter arrangement described in U.S.patent applications Ser. No. 08/130,257 now U.S. Pat. No. 5,434,412 andSer. No. 08/182/,572 now U.S. Pat. No. 5,424,554 in the followingexamples. If this configuration is used to collect light in other thanthe forward direction, the entire lens/aperture/detector assembly wouldbe rotated about an axis centered preferably beneath the surface of theviewed sample.

The optics of the system and methods described herein are such that theyeliminate much of the scattered radiation reaching the detector,specifically all the scattered radiation outside the limited solid angleviewed by the detector, while allowing the use of geometrically wideradiation beams. This enhances the ratio of directly transmitted todiffuse radiation which reaches the detector. This is particularlyadvantageous with a heavily scattering media, e.g., the sample is aportion of an mammalian body, such as a human body. If a human body isused as a sample, preferably a thin region of tissue is selected so asto minimize person-to-person variation in tissue thickness. Once againthe preferred thickness would lie in the range of a few (2-10) mm,comparable to the distances over which the internal light density fallsto a few percent of its maximum value. Thinner tissues can be usedprovided that they are sufficiently vascular to provide good pulsatilesignals. One possible area for use is a finger web which can be clampedto provide a substantially standardized tissue bed thickness. Anotheralternative is the eyelid. Thinner tissues such as these can also bebacked by reflective surfaces to achieve a “double-pass” effect, or withan absorbing backing that also isolates the tissue from underlyingstructures.

The geometrical width of the illumination and detection areas on bothsides of the sample are designed to be wide enough to average overmultiple internal structures such as capillary beds. This reducessensitivity of the results to the exact positioning of the opticalsystem at the chosen site. This approach achieves significantly higherintegrated light intensity at the detectors, while avoiding edge effectsat the extremes of the illumination beam. Thus, the large geometric beamarea facilitates passage of a large number of photons, and the solidangle restrictions limit differences in the lateral spreading of thelight within the tissue at different wavelengths. This geometry producesrelatively higher detected light intensity with more consistent samplingacross tissue inhomogeneities as a function of wavelength than doespreviously described apparatus and methods. The intensity is high enoughthat the fundamental shot-noise detection limit falls below theprecision needed for the trace analytes of interest, and the improvedsampling reduces the non-linearities imposed on the measurementcalibration by scattering.

The preferred working range for the method and apparatus invention isfrom about 700 to about 2500 nm. This region covers the absorbance ofsome of the most preferred substances, including glucose and itsidentifying substances, hemoglobin, deoxyhemoglobin and various drugs ofabuse. The method can be used to determine the hemocrit or to derive theoxygen saturation level in the blood. This method and apparatus can bemade sufficiently rapid to measure arterial pulse data, therebyeliminating another source of potential error. The present improvementscan also be utilized with modulated sources, which are particularlyhelpful in eliminating radiation generated from sources other than whatis transmitted by the sample.

In an other embodiment of the invention, polarized light is used as theilluminating radiation. The detector which collects the radiationincludes an analyzer or filter for transmitting polarized light before,or as part of, the detector, while excluding depolarized lighttransmitted from the sample. Since scattering of polarized radiation bythe scattering sites in the sample will depolarize the scatteredradiation, the use of the polarizer in conjunction with the detectorwill maximize the ratio of polarized radiation to depolarized radiationcollected by the detector. The preferred polarization system has therestricted solid angles of illumination and limited solid angle ofdetection previously described. The reason for this is, in part, thatusing this technique not only eliminates scattered radiation but alsoeliminate radiation from other non-polarized background light sourcesand provides the highest ratio of desired to undesired radiation whileproviding sufficient signal.

If the restricted solid angle geometry is utilized for imaging, an arrayof detector units forming the detector, e.g., a CCD array, can be used.This array, and/or illumination source, can be scanned across the sampleusing any standard mechanical stage or raster scanner, to generate aseries of shadowgrams which can be combined to form a larger shadowgramshowing tissue inhomogeneities. Temperature changes across the tissuecan also show up as an inhomogeneity.

As noted, the present invention also provides a device or apparatus forcarrying out the method of the invention. All the aspects previouslydescribed with respect to the methods can be utilized in the apparatus.

The invention is further illustrated by the detailed description of theinvention and the drawing.

BRIEF DESCRIPTION OF THE DRAWING

FIG. 1 shows several possible characteristics of beams which passthrough a tissue sample;

FIG. 2 shows the preferred embodiment of the invention, a device havinga collimated source, collimation optics on the detector side and fourdetectors arranged for congruent sampling.

FIG. 3A expands the scale of FIG. 2 to illustrate the definition of theangular acceptance ranges by the aperture. FIG. 3B further expands thescale to compare the angular divergences of the present invention withthose of prior art using fiber optics. FIG. 3C shows the angulardivergences typically employed in commercial pulse oximeters with amodification of the present invention shown. FIG. 3D shows theorientation an embodiment to detect reflected radiation from thescattering medium, and FIG. 3E shows two forms of the orientation usedpreviously for the detection of reflected radiation;

FIG. 4 shows two variations on optical systems for the invention, withFIG. 4A having a slightly converging illumination beam while FIG. 4B hasa slightly diverging illumination beam, and the detector on each showingcollimating optics;

FIG. 5 shows an additional embodiment, a device with a converging lensand aperture placed near the focal point of the converging lens;

FIG. 6 shows a system wherein a bifurcated optical bundle may besubstituted for the congruent sampling apparatus shown in FIG. 2;

FIG. 7 shows a path reversed version of the optical system of FIG. 2,showing congruent illumination and collimated beams; and

FIG. 8 shows a polarizer and analyzer system.

DETAILED DESCRIPTION OF THE INVENTION

The present invention is directed, in part, to methods of improvingconsistency of measurement and, therefore, calibration of samples whichhave scattering properties. The methods and apparatus of the inventionare particularly well suited to non-invasive testing of human tissue.The invention is based, to a large extent, on the recognition that ifsufficient signal is generated in a non-invasive test system, limitingthe solid angle such that much of the scattered radiation would beoutside the solid angle of detection would minimize changes in theeffective pathlength caused by unequal scattering in the sample.

One of the advantages described in the Block '265 Patent, as well as theother above-cited United States patent applications, is that the use ofbroad spectrum illumination and detection provides greater signalintensity than in a conventional spectrophotometric systems. Highdetected intensity is vital for the detection of the specific absorptionof trace analytes, because the changes they make in the detected signalsare very small. The signal changes produced by clinical concentrationsof glucose, for example, as so small that they are difficult to separatein aqueous solution from changes due to the displacement of water andslight alterations in solution temperature. As a result, there are noreliable publications of the specific absorption spectrum of glucose insolution in the near infrared. Nevertheless, the scale of glucoseabsorption can be approximated on the assumption that the hydroxylgroups of glucose create an average band whose size and shape aresimilar to the corresponding band of pure water, but whose location isshifted by interactions with adjacent groups on the molecule to have apeak slightly above 1010 nm. At 5.5 mmol/L (100 mg/dL), theseglucose-bound hydroxyl groups are about 2000 times less concentratedthan those of bulk water. By scaling from the 0.2 OD absorption band ofpure water at 960 nm, the shifted band-peak from glucose at 5.5 mmol/Lconcentration would have a magnitude of about 10⁻⁴ OD at 1 cm path,which would create a transmission change of 2.3×10⁻⁴. This result agreeswith the general level observed for this shifted OH band in higherconcentration measurements reported by Koashi in U.S. Pat. No. 4,883,953for the chemically similar saccharose.

To be clinically useful, this glucose concentration should be measuredwith a precision roughly 20 times smaller, which translates to ameasurement resolution or precision at the one part in 100,000 level inthe case when the tissue viewed has an effective pathlength of 1 cm. Ifpulsatile measurement is employed, the time-varying component s of thesignals, which typically comprise 2-10 percent of the total signal, mustalso be measured to sufficient precision. Here, however, the effectivepathlength generating the observed pulsations is much shorter than 1 cm.For example, in the 950 nm range, the dominant absorber in the blood isoxyhemoglobin, with total absorbance of about 3 OD for a midrangehemoglobin concentration. Such an absorbing solution creates 10 percentchange in transmission at a pathlength of only 1/60th cm. As a result,the pulsatile measurement of specific glucose absorption requiresprecision better than one partper-million.

This is significantly below the repeatability achievable with thepresent stateof-the-art in pulse oximetry The photodetector shot noiselimit alone require that there be more than 10¹² photoelectrons detectedin the integration time of the measurement to reach the part-per-millionresolution range. Allowing for the fact that only a portion of eacharterial pulse is useful, the shot noise limit translates to arequirement for intensity levels at the photodetectors which lie in the10-100 microwatt range, after passage through the tissue. With thicktissues and/or inefficient use of intensity, this requirement in turncan lead to a need for 10-100 milliwatts, o r even watts, of radiationlaunched into the tissue. The intensity which can be used has apractical upper limit set, at least, by the levels which becomeuncomfortable for the test subject due to heating by the absorbedradiation.

Sensitivity at the part-per-million range also enables the measurementof the temperature of subsurface features with milli-degree°C.sensitivity, or lower. This follows from the fact that the 960 nm waterband changes its distribution with wavelength when the temperaturevaries, thereby allowing the temperature to be visualized by imagingdetectors just as if it were another absorbing constituent. Ourmeasurements with four 70 nm wide overlapping filters positioned nearthe 960 nm absorption band of water have shown that a 1° C. temperaturedeviation produces signal deviations that are on the same approximatescale as those due to the intrinsic absorption bands of glucose at aconcentration of about 110 mmol/L (2000 mg/dL), but with a differentdistribution between the four detector signals. When the signals fromimaging detectors are analyzed to display the mean subsurfacetemperature along the collimated path through the tissue, the resultant“shadowgram” images allow the localization, for example, of regions ofhigh metabolism, such as tumors.

It is known that light transmitted through tissue or other scatteringmedia can pass through in a variety of ways. FIG. 1 shows a sampling ofthese modes of transmission, in the form of a series of ray traces. Moreparticularly, ray trace A on FIG. 1 shows a “snake like” ray which onlyhas minor, glancing scattering collisions through the tissue. This typeof scattering comprises a forward-scattered ray. Ray trace B shows a“ballistic” ray which travels directly through the tissue without beingscattered or absorbed; that is, it passes directly through the tissue,being transmitted forward as if there were no scattering centers in thesample. Differentiation between “snakelike” and “ballistic” rays isdifficult, and those who attempted time gating normally include bothtypes of rays in their collected data Ray trace C on FIG. 1 is amultiply scattered ray that emerges within the selected angulardistribution range accepted by the detector. This type of ray will bedeemed a forward scattered ray by the apparatus and it will be measuredor included as part of the presumed “unscattered” radiation. Ray trace Dis a multiply scattered ray which scatters outside the selected angulardetection region viewed by the detector. For convenience, this multiplyscattered ray is shown as being only slightly outside the solid angleviewed by the detector but actually multiply scattered rays outside thesolid angle viewed by the detector would constitute the majority of theradiation hitting the sample.

In the visible and near infrared, the relative abundance of thesevarious ray types is a very strong function of the thickness of thetissue, the magnitudes of the scattering and absorption cross sectionsfor the selected tissue sites, and on the socalled single scatteringphase function (sometimes called the angular scattering distribution).The cross sections and phase functions vary strongly between tissuetypes, and within given tissues depend strongly on the state ofhydration and patient-dependent mixture of physiological sub-structures.

There has been much attention paid recently in the literature on methodsto measure, interpret, and extrapolate from these fundamentalmeasurements to obtain single scattering parameters and bulk absorptionscattering coefficients for the media. The theoretical problem, however,is complex and does not lead to closed-form solutions unless severelyrestrictive assumptions are made. The situation is further complicatedby the fact that the coefficients themselves are strongly tied to theunderlying theoretical model, so that much care must be taken even tocompare coefficients obtained by different workers to be sure that theyare defined and measured in the same way. In general, the theoreticalmodels have not yet been able to deal well with nonhomogeneous tissue,and provide only general guidance on the scale and functionaldependencies to be expected in real measurements.

As a result of the complexity of theoretical models, the Monte Carlocalculational method is often employed. Here, the paths of many photonssuch as those indicated in FIG. 1 are followed through their transitacross the scattering medium. The frequency of interaction and the angleof each scattering are selected randomly by the computer so as to beconsistent, on the average of all interactions, with the assumedfundamental cross sections and phase functions which form the input datafor the calculation. While this method readily accommodatesnonhomogeneous media, it is very ponderous for the solution of theinverse scattering problem in which the cross sections and phasefunctions are determined from the observed experimental measurements.Here again, the results provide only general guidance as to how themeasurements will vary with experimentally controllable parameters.

There is agreement in the literature, however, that in the limit ofsmall angular acceptance for a parallel initial beam incident on asemi-infinite plane slab of tissue or other scattering media, thetransmission of light in the forward direction follows the classic formof Beer's law:

T=e−(μ_(s)+μ_(a))x  [1]

where μ_(s) designates the coefficient of scattering, μ_(a) is the totalabsorption coefficient made up of the sum μ_(a l=μ) _(a1)+μ_(a2)+μ_(a3)+. . . of the absorption coefficients of the individual constituentspresent in the scattering media, and x is the thickness of the slab.While μ_(a) is relatively independent of the experimental setup, μ_(s)depends strongly on how small a solid angle the detector subtends in theforward direction (i.e., the smaller the angle, the easier it is for ascattering event to remove a photon from the detected beam, and thelarger μ_(s) becomes). When the detector solid angle and/or thethickness are small enough that equation [1] is followed, scatteringaffects the transmission with the identical functional dependence as anyof the absorbing constituents. Further, the contributions of layeredmedia can be readily calculated by dividing x into a sum over theindividual layer thicknesses. Note that with this exponential dependencethe effect of a strongly absorbing layer on the transmission isindependent of the layer depth within the composite medium; such a layerproduces the same percentage change in the transmission wherever it liesalong the optical path.

There are many different theoretical models in use to describe thetransport of light across scattering media, most of which are based onthe available closed form solutions to a diffusion equation. To couplethese available solutions to the real scattering problems, these modelstypically assume a point source of light, completely diffuseillumination, or a rapid transition from a parallel input beam of lightto a highly diffused beam. Once such diffuse illumination isestablished, it typically propagates through the tissue with arelatively stable angular distribution that rapidly approaches anisotropic one whose magnitude decays with distance following afunctional form similar to:

T=e−{square root over (3+L μ′_(a)+L (μ′_(s)+L +μ′_(a)+L ))} x  [2]

Here the prime (′) indicates that the coefficients in equations [2] arenot the same as in equation [1]. The magnitudes of the coefficients arein fact quite model dependent, while the functional form is generallymodel invariant.

Equation [2] is only applicable for one dimensional models in which theangular distribution of the light reaching the detector is ignored.Theoretical models with three dimensional geometries produce equationssimilar to equation [2], with T replaced by the photon density, xreplaced by the radius r from the coordinate origin, and with additionalfactors of 1/r² present to account for the loss in absolute photondensity as the approximately isotropic beam spreads radially.

A similar form also develops in the well-known Kubelka-Munk scatteringtheory, with the characteristic exponential decay constant 3μ_(a)′(μ_(s)′+μ_(a)′) of equation [2] replaced in more complicatedcombinations of exponentials by K(2S+K), where K and S represent bulkabsorption and scattering coefficients of the media defined specificallyunder the constraint of perfectly diffused incident and transmitted (orreflected) light distributions. The slight difference in the functionalform of these exponential terms results from the details of thedifferential equation whose closed form solutions are used toapproximate reality. Thus, in the Kubelka-Munk formulation a onedimensional differential equation is assumed with cross-couplingconstants K and S between two isotropically distributed light beamsprogressing in the forward and backward directions. All the threedimensional information is carried in the assumption ofwell-maintained-isotropy. Despite their differential genesis, it isclear that suitable renormalization can carry the K-S form into theμ_(a)′−μ_(s)′ form.

Over most of the 700-1400 nm range, μ_(a)′ is much smaller than μ_(s)′(see Wilson et al., IEEE J Quant Elec 1990; 26, 2186-99 for a review),and the second term under the square root in equation [2] is oftendropped in the theoretical analyses. Above 1400 nm, μ_(a)′ undergoes atenfold increase because of increased absorption by water. The exponentof the transmission in equation [1] is often referred to as the“transport mean free path (mfp),” with μ_(s)′ also including acorrection for the average cosine of the scattering phase function. Fortypical tissues, the mfp usually lies in the range of 0.05-0.2 mm.Similarly the exponent of the transmission in equation [2] is called thepenetration depth, i.e., the distance at which the diffuse radiationlevels fall to 1/e. For typical combinations of coefficients and phasefunctions, this is 10 to 20 times larger than the mfp, and ranges up toa few mm.

Monte Carlo models, such as those of Flock et al., IEEE Trans Biomed Eng1989; 36, 1162-8, provide some guidance on the relative size ofcollimated and diffuse radiation beams. FIG. 6 in the Flock et al. paperis a polar plot of intensities for “ten bins of equal solid angle.”While the most forward bin shown in this figure appears to have asmaller solid angle than the others, the figure does indicate that thecollimated and diffusely transmitted photons will have roughly equalintensity at depths of about 20 transport mean free paths, comparable tothe natural penetration depth of the diffuse radiation. The distance forsuch equivalence is clearly a strong function of the angles andfundamental coefficients and those of the site.

The important result is that the scattering and absorption effects aremixed together, so that neither acts alone; an increase in one of themwith wavelength or sampling site enhances the impact of the other on thetransmission. When scattering dominates over absorption the dependenceof the exponent on μ_(a) in equation [2] changes from a linear onesimilar to equation [1] to a square root dependence, and the usuallogarithmic transformation of Beer's law that linearizes equation [1]can no longer linearize equation [2]. It is worth noting that equation[2] does not converge smoothly to equation [1] as the scatteringcoefficient goes to zero because of the extra factor of 3 in theequation [2] exponent, again highlighting the model dependentnormalization differences.

The difficulties inherent in the calibration of diffuse transmission areillustrated by the work of Cope et al., SPIE 1431, 251-62 (1991), whocoalesce the square root and 3μ_(s)′ factors in the exponent of equation[2] into a new variable they call the differential pathlength factor(“DPF”).The DPF is a locally linear approximation to the slope of thesquare root dependence after logarithmic transformation which must becalibrated separately at each wavelength of interest with the implicitassumption that the bulk coefficients μ_(a)′ and μ_(s)′ and thethickness x′ will remain constant between the calibration and analysissamples. Cope et al. found that use of such a wavelength dependent DPFwith these assumptions “significantly improved the residuals generatedby multilinear regression analysis.” At the same time, they also notethat the calibration of the DPF is impractical in vivo, where μ_(a)′ cannot be varied independently, and instead propose an additionalmeasurement of the mean time of flight of detected photons as a measureof the DPF. Their calibration method is clearly highly complex even forthis case in which the analyte itself contributes roughly 1/10 of thetotal bulk absorption coefficient μ_(a)′. This is 1,000-10,000 timehigher than the equivalent relative contribution of glucose to the totalabsorption in the present invention.

As noted in the U.S. patent application Ser. No. 08/333,758, the smallsize of the analyte absorption allows a convenient alternativecalibration method in which the effect of the analyte is to create smalllinear perturbations of the signal intensities from a set of referenceintensities which are defined by the absorption of major constituents,such as the hemoglobins, and the thickness and scatteringcharacteristics of the sampled site. These reference intensities arethemselves highly non-linear, comprising a mixture of the functionaldependencies of equations [1] and [2]. The preferred methods describedin the Block '265 Patent and the related application take advantage ofthe broad and shallow wavelength variations of the major absorbers andthe scattering characteristics of the tissue to facilitate the accuratedetermination of the reference intensities, and do so dynamically oneach individual measurement so as to adjust automatically for theinevitable physiological changes in the selected sites from day-to-day.The use of restricted solid angle, particularly in the form ofcollimation carries the functional behavior closer to that of equation[1] as the effects of scattering, absorption, and thickness becomedecoupled so that the estimation of the reference intensities becomemore reliable and more easily calibrated. This trend applies across thewhole spectrum of calibration methodology disclosed in the relatedpatent and applications, including the simultaneous intercomparison ofmeasurements on multiple sites and/or with multiple detectorconfigurations with different wavelength responses.

Most importantly, neither the theoretical models outlined above nor theexperimental measurement systems employed using classical methods dealwell with the inhomogeneities in tissue and other layered scatteringmedia. Closed-form solutions only appear for extremely limitedassumptions of the mutual variation of μ_(a)′ and μ_(s)′ with depth intothe scattering medium. The Monte Carlo approach can handle tissueinhomogeneities somewhat better, but also require input values for thecoefficients and the angular distributions of single scattering eventsthat are very difficult to obtain reliably. Even the early experimentalwork of Kubelka, J Opt Soc America (1954) 44, 330-5, demonstrated thatin diffuse light, the impact of an inhomogeneous layer is very stronglydependent on its depth when viewed in reflectance, and somewhat lessstrongly when viewed in transmission. His transmission results alsoshowed a remarkable symmetry of the effect of a strongly absorbing layerin which it has the identical impact on the total transmission in twolocations that are symmetrically offset from the midpoint of the medium.

Little theoretical work has been done as yet on the impact of lateralinhomogeneities within individual layers in the scattering medium. Theseare on the one hand equivalent to classical wedging errors inspectrophotometry, and on the other hand potentially more complicatedbecause of the non-linearities inherent in the admixture of diffuselight. At the same time, the available sites for non-invasivemeasurement of glucose and other trace constituents can not be expectedto be homogeneous at the part-per-million level within a givenindividual, let alone across a population of individuals.

One preferred embodiment of the present invention, which is designed toaddress the simultaneous need for high detected intensity and toleranceof inhomogeneity in the sampled tissue, is shown in FIG. 2. This opticalsystem represents an improvement over those taught in the '265 Patentand related applications, in that additional optical elements have beenadded to decrease the angular range of light leaving the tissue whichwill be accepted into the detector elements. The attendant loss ofintensity at the detectors is compensated by increased size of thedetectors, beamsplitters, and the geometrical area of the incident anddetected light beams.

This optical system employs collimating optics for both illumination anddetection, with the detector having a plurality of detector units placedsuch that they achieve congruent sampling. Radiation source 10 isselected so that it provides broad spectrum illumination, e.g., 700-2500nm illumination. Radiation from radiation source 10 passed throughcollimating lens 12 before striking tissue 20. Optional aperture 14 isshown which helps define the collimation optics in conjunction withcollimating lens 12. Tissue sample 20 may be any sample but a thin,repeatable sample such as a finger web is preferred. For this type ofthin sample, the amount of scattering is minimized and it is possible tohold the web and compress it such that a standardized thickness may beachieved. The more standardized the thickness is, the more likelyuniversal calibration (or calibration to a limited group of thicknesses)can be. If universal calibration can be achieved, the calibration canset at the factory and corrections for effective pathlength can be madein the instrument itself. If such universal calibration is not achieved,some calibration measurements may be required before meaningful data canbe obtained.

Once the radiation has traversed tissue sample 20 and exits the tissuethrough area defining aperture 25, it passes through detectorcollimating optics 30 formed of converging lens 32, aperture 34 andrecollimating lens 36. This type of collimating optics is conventionallyused in telescopes and other devices where collimation of light isdesired. The collimated beam exiting collimation optics 30, specificallyrecollimating lens 36, then goes through a series of beam splitters 42A,42B and 42C and onto four detector units 44A, 44B, 44C and 44D. The beamsplitters and detector units are arranged such that the entire detectionunit 40 provides congruent sampling of the beam. More particularly, thebeam splitters and detector units are arranged such that the pathlengthand angles from recollimating lens 36 to any of detector units 44 areequal and each of detector units 44 are optically superimposable uponthe other. More details of this type of congruent sampling is set forthin U.S. patent application Ser. No. 08/130,257. Although four detectorunits are shown, the exact number may be varied.

FIG. 2 also shows an additional testing or detector unit 50 (shown as a“black box”) which is not in line with the collimation or anglerestricting optics. In some circumstances, the scattered radiation mayprovide information in addition to or supplementing that obtained fromthe unscattered radiation. Additional detector unit 50, which mayactually be a plurality of detector units, can be used at differentpositions and angles, thereby testing the scattered radiation andproviding additional information. One particularly valuable embodimentuses this detector unit in reflectance or transflectance mode, that ison the same side of the tissue as the illuminating optics. Thisadditional detection unit can be used with any of the embodiments shownherein.

The optical system of FIG. 2 preserves spatial information, in thatthere is a one-to-one correspondence between a location on the viewedtissue surface area and a location on the active area of each detector.With adjustment of lens-aperture-lens-detector distances, a reasonablysharp focus can be achieved, particularly as the accepted viewing solidangle shrinks towards perfect collimation. As a result, replacement ofthe detectors shown in FIG. 2 by imaging devices such as CCD's or otherarray detectors creates an imaging system that produces “shadowgrams”using “ballistic” or “snakelike” photons, with the added advantage ofsimultaneous congruent imaging in all pixels in the detector arrays. Thesimultaneity facilitates the real-time processing of the signals to formtuned images of subsurface structures in the different analytes,including temperature. Since certain tumors are known to have adifferent temperature than surrounding tissue, this imaging system hasuses for detection of tumors and other anomalies. The illuminationsource and detector array can be scanned across a large tissue sampleand combined to form a larger shadowgram.

The essence of the present invention relative to the prior art isillustrated in the sequence of FIGS. 3A-E. FIG. 3A is an expanded scaleversion of FIG. 2, showing the way in which the collecting lens andangle defining aperture function to limit the acceptance cones of thelight emerging from the tissue at different point on the exit area.Light emerging from the tissue at angles outside the acceptance conefrom anywhere in that exit area, defined by aperture 25, strike theangle defining aperture outside its central opening. As the relativemagnitudes of the forward collimated light intensity and the diffusescattered light intensity are not know exactly from either theory orexperiment, as explained above, the aperture size and the resultantangular acceptance cone must be tailored to the particular observationsites selected.

FIG. 3B illustrates the difference between the present invention (top)and prior art (bottom) systems using fiber optics to introduce incidentlight to the scattering media, and then transport the transmitted lightto analysis means. The angular ranges over which incident and detectedlight are launched or received by the optics are indicated by thearrows. These angular range for fiber optics cover their full numericalaperture, which typically comprises an ˜50 degree full angle cone, whilethe present invention illuminates the tissue with nearly parallel light,and detects light emerging in only a small angle cone whose full angleis less than 20 degrees, and preferably less than 10 degrees. Lightentering along any of the rays shown in the figure will spread sidewaysabout that ray due to scattering, and the amount of spread will varywith the wavelength of the light. Light which has spread laterally has amuch higher probability of reaching the detector in the prior artarrangement shown because the receiving fiber optic will accept lightover a much higher solid angle. The arrangement shown for the presentinvention configuration provides benefit even for the detection ofdiffuse radiation because the contributions of inhomogeneities near theedges of the observed areas will be more evenly sampled at allwavelengths. In addition, to the extent that the acceptance cones,tissue thickness, and tissue scattering parameters can be chosen toinclude predominantly unscattered collimated light reaching thedetectors, the variation of that detected light with analyteconcentration will trend from the inherently complex form in equation[2] to the inherently more readily calibratable form of equation [1].

FIG. 3C illustrates the optical geometry employed in most commercialpulse oximeters. Here, LEDs (50) at the selected wavelengths are placedadjacent to each other on one side of the tissue, and emit theirradiation into the tissue at angles similar to those of the previouslydescribed fiber optics. The detector is placed on the far side of themonitored tissue, with an acceptance cone for light leaving the tissuethat can approach the full hemisphere. Again, scattering spreads thelight sideways about each incident ray, and light which the detectorsreceive will have sampled a region much wider than the spacing betweenthe two detectors. This light includes contributions from tissueinhomogeneities which lie at the outer edges of the indicated rays, andsince the degree of spreading is wavelength dependent, theseinhomogeneities make different contributions to the signal at eachwavelength. This effect is a major contributor to the high sensitivityof pulse oximetry results to motion of the site, compression of the siteby the instrument, and to small changes in tissue thickness, all ofwhich are greatly reduced in the present invention. To meet the presentinvention, detector 54 is broadened in area as shown by the dotted linesto be at least comparable in size to the sample thickness.

FIG. 3D illustrate the way in which the optical system of the presentinvention can be employed to detect radiation reflected or backscatteredfrom the monitored site. This arrangement is particularly advantageousfor sites such as the forehead or eyelid, where it is impractical toplace a detector on the far side of the tissue. Here, the incident andemerging light beams preferably cross within the tissue. Once again, theadvantage of this optical arrangement, even for detection of diffuse asopposed to singly scattered radiation, is that the relative contributionof tissue inhomogeneities near the edge of the illuminated area arerendered more equal between detectors at different wavelengths. Theparticular angles shown should be taken as illustrative rather thanrestrictive; there are numerous alternative arrangements well known inthe art which provide additional advantages. One worthy of mention isthe case in which the incident radiation would enter normal to thesurface, with the restricted angle acceptance cones lying along (all or)part of an annulus centered about the incident beam.

FIG. 3E illustrates two optical configuration often employed in theprior art for the measurement of reflected light. The fiber optic(60/62) configuration shown suffers from the same limitations as that inFIG. 3B, with the wavelength dependent differences in scattering andhence penetration to deep inhomogeneities skewing the relative signalbetween detectors even more strongly. The 2nd configuration in FIG. 3E,with LEDs (70) and the detector (74) on the same side of the monitoredsite, has still greater potential for unequal sampling of both lateraland axial inhomogeneities because of the larger angular acceptance ofthe detector.

Although FIG. 2 shows a collimated beam for illuminating the tissue, insome circumstances it may be better for the illumination beam to be notperfectly collimated FIG. 4 shows two variations which implement such anon-collimated beam condition; FIG. 4A shows a converging beam fortissue illumination while FIG. 4B shows a slightly diverging beam fortissue illumination. More particularly, the apparatus of FIG. 4A has aradiation source 110 which generates a radiation beam that passesthrough a converging lens 112 before striking tissue 120. The focalpoint of converging lens 112 is on the opposite side of tissue 120 fromradiation source 110. The radiation transmitted through tissue 120passes through converging lens 132, preferably an aperture 134, and arecollimating lens 136.

FIG. 4B shows substantially the same system as FIG. 4A except lens 212is a slightly diverging lens as a compared with the slightly converginglens 112 in FIG. 4A. Radiation source 210, lens 232, aperture 234 andlens 236 are the substantial equivalent of their corresponding numberedparts (110, 132, 134 and 136, respectively) in FIG. 4A. While FIGS. 4Aand 4B illustrate an output beam from the optical system which iscollimated, it may be that not all of optics 130 in FIG. 4A or optics230 in FIG. 4B is necessary since a slight variation from collimation onthe output beam may be desirable to get the best ratio ofsnakelike/ballistic to scattered rays. The configuration chosen is theone which empirically optimizes the transmitted photon intensity whilemaintaining insensitivity to internal inhomogeneities in the tissue.

FIG. 5 shows another variation on the optical system of the invention,one with a more highly converging lens 312 which has a focal point onthe far side of tissue 320. An aperture 334 is placed at or near thefocal point of lens 312 and a recollimating lens 336 is placed nearaperture 334. By placing aperture 334 at the focal point of lens 312, apinhole camera-type system is created whereby the image of tissue 320 isreversed but formed directly upon the detector. Again, this opticalsystem may have advantageous properties depending on the type of tissueor other sample measured.

FIG. 6 shows a variation of FIG. 2, whereby instead of the beam splitterapparatus 40 shown in FIG. 2, a bifurcated optical bundle which is splitinto four parts, each leading to a different detector unit, issubstituted. If the detector units are located such that the length ofthe optical fiber leading to the particular detector unit is identical,this system provides an approximation of the congruent sampling shown inFIG. 2. More details concerning this type of bifurcated optical bundleis described in U.S. patent application Ser. No. 08/130,257.

FIG. 7 shows a system using the beam splitter array of FIG. 2 reversedfor congruent illumination rather than congruent sampling. Fourradiation sources 710A, 710B, 710C and 710D, are used to illuminate thetissue sample. The radiation issuing from each of the radiation sourcesgoes through a collimating lens (712A, 712B, 712C and 712D,respectively) and then is redirected by one of the beam splitters 716A,716B or 716C to illuminate tissue 720. The radiation transmitted bytissue 720 passes through converging lens 732 and aperture 734 beforestriking detector 744. Optionally, an additional lens 736 (not shown)could be used to recollimate the transmitted radiation before it strikesdetector 744. The radiation sources, collimating lenses and beamsplitters are arranged to provide congruent illumination and eachseparate radiation source may have an associated modulator to provide adifferent modulation to the radiation issuing from that radiationsource. This type of modulation apparatus, and its advantages, isdescribed in more detail in U.S. patent application Ser. No. 08/182,572.Briefly, using a plurality of modulators each providing a differentmodulation to the associated radiation issuing therefrom, and using aform of modulation differentiation at the detector (such aselectronically separating the signals based on modulation frequency)provides a method which allows differentiation at the detector of thesource of the illuminating radiation, and accordingly allows additionalinformation to be generated from a single detector. For example, if theradiation sources cover different wavelengths, a single detector candifferentiate the intensity of the transmitted radiation at eachwavelength range by using the modulation to determine the wavelengthrange. This can eliminate the requirement of the system illustratedabove which requires a plurality of detector units. For improvedresults, both the congruent illumination shown in FIG. 7 and thecongruent sampling shown in FIG. 2 may be used in the same device.

FIG. 8 shows a different embodiment of the invention whereby a polarizerand an analyzer are used in conjunction with the optical system in theinvention to provide signal differentiation. More particularly,radiation source 810 emits radiation (shown with the polarizationdistribution 812) which then passes through polarizer 815. Onlypolarized radiation (see 816) illuminates tissue 820 and since the actof diffuse scattering in tissue 820 will depolarize the scattered light,only unscattered (or forward scattered) radiation is transmitted fromtissue 820 as polarized radiation. The radiation is then transmittedthrough analyzer 850, which is a polarizer that passes only radiationwhich has the same polarization as polarizer 815, and detector 844 (notshown) detects this polarized radiation to yield a signal.

One advantage of using this polarizer/analyzer system is that itimmediately segregates scattered from unscattered radiation, sincescattered radiation is depolarized and cannot pass through analyzer 850to detector 844. Accordingly, although collimation optics and therestricted solid angle can be used with the apparatus of FIG. 8, theseare not absolutely necessary because the polarizer/analyzer pair willeliminate the off-angle scattered radiation in any case.

In all of the foregoing embodiments, there is an advantage to usinggeometrically broad beam radiation as opposed to narrow beam radiation.The use of such broad beam radiation provides a higher input signalwhich, when constrained by the solid angle restrictions of the presentapparatus, still provides sufficient illumination and signal to meet theprecision requirements for the analysis of trace constituents. Further,once the signals have been received, the processing described in theBlock '265 Patent and the previously cited applications may be appliedto differentiate signal from background and obtain meaningful data.

The advantages of the present invention apply to spectrophotometricsystems such as those employed in pulse oximetry. While the shot-noiseconstraints on the detected intensity are lower because the absorptionof the hemoglobins are so much larger the acceptance angle restrictionsprovide greater linearity and improved calibratability, as well asreduction in the severity of motion and breathing artifacts, and otherlimitations on universality of calibration.

For the analysis of trace constituents where the high photon fluxrequirement is critical, the present invention is particularlyadvantageous when combined with the use of broadband and broadbandoverlapping detectors, as taught in the Block '265 Patent and the parentapplications.

The foregoing description is meant to be explanatory only and is notintended to be limiting as to the scope of the invention. The inventionis defined by the following claims.

What is claimed is:
 1. A device for imaging inhomogeneities in a tissuecomprising: a radiation source for illuminating a designated area ofsaid tissue with radiation having a wavelength between 700-2500 nm, saiddesignated area being at least comparable in size to the thickness ofsaid tissue, said illumination of said designated area being over arestricted solid angle; a detector comprising a plurality of detectorunits in an array, said detector arranged to receive radiation from onlya limited solid angle with an area comparable in size to said designatedarea; whereby a shadowgram of said inhomogeneities in said tissue isgenerated.
 2. The device of claim 1 further comprising scanning meansfor scanning said radiation source and/or said detector across saidtissue.
 3. The device of claim 1 wherein said inhomogeneities comprisetemperature inhomogeneities.
 4. The device of claim 1 wherein saidinhomogeneities comprise concentration inhomogeneities.
 5. The device ofclaim 1 wherein said inhomogeneities comprise scatteringinhomogeneities.
 6. A method of imaging inhomogeneities in a tissuecomprising: illuminating a designated area of said tissue with aradiation source generating radiation having a wavelength between700-2500 nm, said designated area being at least comparable in size tothe thickness of said tissue, said illumination of said designated areabeing over a restricted solid angle; detecting transmitted radiationwith a detector comprising a plurality of detector units in an array,said detector arranged to receive radiation from only a limited solidangle with an area comparable in size to said designated area; whereby ashadowgram of said inhomogeneities in said tissue is generated.
 7. Themethod of claim 6 further comprising the step of using a scanning deviceto scan said radiation source and/or said detector across said tissue.8. The method of claim 6 comprising the step of measuring temperatureinhomogeneities by comparing said detected inhomogeneities with atemperature standard.
 9. The method of claim 6 comprising the step ofmeasuring concentration inhomogeneities by comparing said detectedinhomogeneities with a concentration standard.
 10. The method of claim 6comprising the step of measuring scattering inhomogeneities by detectinginhomogeneities at various angles.